The invention relates to radiation detectors in nuclear medicine and more specifically to the electronics used to detect signals from photodiodes used in some radiation detectors.
Nuclear medicine imaging assesses the radionuclide distribution within a patient after the in vivo administration of radiopharmaceuticals. The imaging systems that asses the radionuclide distribution comprise radiation detectors and associated electronics. The imaging systems detect x-ray or gamma ray photons derived from the administered radionuclides. Single photon emission imaging and coincidence imaging are two forms of nuclear medicine imaging that are currently in common use. In single photon emission imaging, the radionuclide itself directly emits the radiation to be assessed. For example, in Single Photon Emission Computed Tomography (SPECT), γ-emitting radionuclides such as 99Tc, 123I, 67GA and 111In may be part of the administered radiopharmaceutical. The imaging system may use a lead collimator to eliminate all photons but those photons perpendicular to the surface of the detector. The location and energy of emitted photons may then be accumulated until a satisfactory image is obtained. Coincidence imaging eliminates the need for such a collimator by relying on the detection of two photons at different detectors at nearly the same time. An example of coincidence imaging in current clinical use is Positron Emission Tomography (PET). In PET, β+-emitting radionuclides such as 11C, 13N, 15O, 18F, 68Ga, 82Rb are part of the administered radiopharmaceutical. The emitted positrons react with electrons within the patient's body, the annihilation creating two 511 keV photons emitted in opposite directions. The two photons are then detected within a certain time window, generally in the nanosecond range, of each other. Radiation detectors for nuclear medicine imaging may need to detect photons from 1 keV to several MeV in energy. Solid state detectors such as those made from Cadmium-Zinc-Telluride that convert photons of such energy directly to electrical signals are now on the market. However, solid state detectors have issues such reliability, longevity, resolution, and cost. Moreover, such systems must be cooled, adding to cost and reliability concerns.
Therefore, many radiation detectors in current use in nuclear medicine imaging systems consist of a scintillation crystal, or scintillator, for converting x-ray or gamma ray photons into visible light photons, so called scintillation photons, and a device for converting the scintillation photons into electrical signals. The Anger camera pioneered this approach in the 1950s is and more fully described in U.S. Pat. No. 3,011,057. The Anger camera consists of a NaI crystal and an array of photomultiplier tubes (PMTs). In operation, gamma ray photons cause scintillation events in the Nal crystal. The resulting scintillator photons then impinge the different PMTs. The different signals amplified by the different PMTs yield information about the location of the scintillation event within the NaI crystal.
In the current art, a variety of scintillation crystals may be used, depending on how the crystal is to be used (e.g., SPECT or PET), cost, reliability, resolution and speed of imaging. Crystal material includes sodium iodide (NaI), cesium iodide (CsI), barium germanate (BGO), barium fluoride (BaF2), lutetium oxyorthosilicate (LSO), and others.
Regardless of the scintillator, the scintillation photons produced must be converted into an electrical signal to be analyzed. PMTs are still often used. A PMT is a vacuum tube including a photocathode, and an electron multiplier sealed into an evacuated glass tube, and an input window which is optically coupled to the scintillation crystal. Scintillation photons (4 or 5) incident on the photocathode cause the photocathode to emit an electron. The electron is absorbed by a dynode which emits 5-6 electrons. A series of dynodes repeat this reaction until a final large cluster of electrons is fed through the anode as a pulse to the attached logic circuits to determine position.
PMTs are extremely sensitive to low levels of light. However, PMTs have many drawbacks. PMTs require a high voltage for operation, typically greater than 1000V. PMTs are vulnerable to drifting in performance, especially early in their life cycle. PMTs are susceptible to mechanical failure and may thus be less reliable. PMTs are susceptible to magnetic fields, such as from the MRI devices (and even from the earth's comparatively weak magnetic field). PMTs are physically bulky. The size of the PMTs determines and limits the intrinsic spatial resolution of a detector system. Furthermore, PMTs require lead shielding, thus increasing the weight of the overall camera. This increases costs, especially in the case the camera must be moved by motors for tomographic imaging.
In addressing the above problems, photodetectors composed of an array of solid state photodiodes have been used rather than PMTs. See, for example, U.S. Pat. No. 5,171,998. Inorganic photodiodes, generally various forms or compounds of silicon, address some of the problems of the PMTs. The inorganic photodiodes are more stable over their life cycle, more robust mechanically, not susceptible to magnetic fields, and much smaller and lighter. However, inorganic photodiodes have their own disadvantages. They are expensive, difficult, and slow to fabricate. Their mechanical structure is rigid. Inorganic photodiodes are susceptible to radiation damage. Inorganic photodiodes generally have a poor spectral response to long wavelength scintillation photons from certain scintillation crystals, such as CsI. Finally, the low band gap of silicon based photodiodes yields thermally generated leakage current, which acts as noise in associated circuits which read the signal from the inorganic photodiodes. The silicon photodiodes must be cooled to lower such leakage current to acceptable levels.
The use of carbon-based photodiodes in lieu of inorganic photodiodes has been disclosed in co-pending, commonly assigned U.S. patent application Ser. No. 10/369,944 filed Feb. 19, 2003, entitled Carbon-Based Photodiode Detector for Nuclear Medicine by Brabec et al. The many advantages of using an array of carbon-based photodiodes in a photodetector are discussed in the co-pending application. One such advantage is a substantially lower leakage current (or “dark” current) than found in current inorganic photodiodes. Such substantially lower leakage current is due to the higher band gap seen in carbon-based photodiodes. However, the capacitance of such photodiodes is substantially increased.
Current front-end electronics coupled to radiation detectors using silicon photodiodes have been optimized to minimize noise and thus maximize energy resolution. However, the application of the same front-end electronics to radiation detectors using carbon-based photodiodes fails to yield substantial improvements in energy yield due to the reduced noise from the leakage current. Thus their remains a need in the radiation detector art for front-end electronics that take advantage of the different electrical characteristics of low leakage current photodiodes, and carbon-based photodiodes in particular.
In accordance with an embodiment of the present invention, a method of selecting an integration time is used for a number of shaper circuits of a radiation detector system having a scintillator optically coupled to an array of low leakage current photodiodes. The lower and upper limit on the range of integration times is determined, and then selecting the integration time within the range of integration times without further regard to the energy resolution to the system.
In accordance with an embodiment of the present invention, a radiation detector having a scintillator, an array of low leakage noise photodiodes optically coupled to the scintillator, and a number of shaper circuits. Each shaper circuit is electrically coupled to one photodiode of the array of photodiodes. Each of the number of shaper circuits has an integration time that is within a range of integration times within which energy resolution is not a consideration.
In accordance with an embodiment of the present invention, a low noise radiation detector for generating a signal having both parallel and serial noise triggered by an x-ray or gamma ray events. The radiation detector includes a photodetector having an array of carbon-based photodiodes and a shaper circuit electrically coupled to the photodetector. The shaper circuit has an integration time wherein which is greater than the integration time of another shaper circuit of another radiation detector having another photodiode including an array of silicon-based photodiodes.
In accordance with an embodiment of the present invention, a low noise radiation detector system having a gantry and a radiation detector head mounted on the gantry. The radiation detector includes a scintillator and a photodetector having an array of carbon-based photodiodes. The radiation detector further includes associated electronics such as a shaper circuit electrically coupled to the photodetector. The system also includes a computer in communication with the radiation detector. The integration time of the shaper circuit is greater than the integration time of another shaper circuit of another radiation detector having another photodiode including an array of silicon-based photodiodes.
In accordance with an embodiment of the present invention, a method of selecting a range for the integration time of a shaper circuit of a radiation detector system having low leakage current photodiodes, The method includes determining the ballistic integration time which is the integration time at which the ballistic deficit has negligible effect and determining the noise integration time which is the integration time at which the parallel noise and series noise are approximately equal. Then the method includes selecting a lower limit of the integration time as the greater of the ballistic and noise integration time and determining an upper limit of the integration time. The upper limit of the integration time is not derived from the parallel noise of the radiation detector system.